The present invention is of use in the field of imaging devices for use in medical diagnosis as well as in the field of digital data acquisition.
The present invention relates to a data acquisition circuit, and more particularly to a data correction circuit for use in positron ECT.
A positron emitted from a positron emitter combines with an electron and is annihilated. The results is that two gamma rays, each of which acquires 0.511 MeV of energy, are discharged in directions nearly making an angle of 180.degree. with each other.
In positron ECT, therefore, two gamma rays detected at a time are regarded as indicative of a positron emitted and annihilated. For the purpose of such detection, it is most common to connect a coincidence circuit to each pair of detectors which face to each other. Every time a coincidence circuit is actuated, an additional 1 which indicates that an event has occurred is written into a particular location in a memory allotted to the pair of detectors connected to the actuated coincidence circuit. Consequently, the readout from this particular location after the end of the data acquisition process indicates the number of positrons emitted and annihilated along the line extending from one of the abovementioned pair of detectors to the other.
FIG. 3-(1) shows the construction of an ordinary memory matrix used in the above-described data acquisition. This memory matrix has an address capability of the number of detectors falling under one set multiplied by the number of detectors falling under the other set. In FIG. 3-(1), the ordinates denote the numbers i given to the detectors falling under one set, and the abscissas denote the numbers j given to the detectors falling under the other set.
A point (i,j) in this orthogonal coordinate system indicates that one of two gamma rays, which have been discharged as a result of the annihilation of a positron, has been incident on a detector i falling under one set, while the other of the two gamma rays has been incidendt on a detector j falling under the other set. Because (i,j) and (j,i) indecate one and the same pair of detectors, the accumulation of data occurs only in the region provided with oblique lines in FIG. 3-(1).
It is most common to reconstruct the image by convolution and back projection. Data to be subjected to these algorithms, i.e., a plurality of points (i,j) in FIG. 3-(1), are collected in a straight direction, and this direction is everchanging. These data have to be rearranged for the reconstruction of the image so that an angle .theta. (FIG. 4) made with the positive direction of a reference axis by a line perpendicular to the above-mentioned straight direction may be plotted on an ordinate against the distance of the straight direction from the origin of the reference axis on an abscissa. The data plotted in this manner are called sinogram, which is shown in FIG. 3-(2). These data are subjected to convolution by means of a reconstruction filter and then to back projection so as to be reconstructed in to a tomographic image.
So far, high resolution has been attained by increasing the number of detectors, with the result that the address capability of a memory matrix has been enlarged from 64.times.64 to 128.times.128, then to 256.times.256, and then to 512.times.512. Consequently, more time has come to be required for the above-described processing from raw data to a sinogram, and the final image formation is apt to be prevented from taking place in real-time.
Especially when an object to be imaged is in motion, dead-time correction has to be carried out on the basis of mean values, by which a large error is introduced. This holds true for the decay correction as well.
Another troubles is that the conventional positron ECT requires software for allowing the processing to proceed from raw data shown in FIG. 3-(1) to a sinogram shown in FIG. 3-(2).